This application relates to apparatus for and methods of performing cryosurgery and in particular to the prevention of damage to tissue surrounding that which is being treated.
Cryosurgery is a modality utilising freezing for ablation of unwanted tissue—e.g. that of a tumour. Clinical experience indicates that for the best results and to ensure proper ablation of the necessary tissue, the tissue should be cooled below about −40° C. The effectiveness may be further enhanced by thereafter allowing the tissue to thaw and repeating the freezing process.
Cryosurgery has great potential for effective treatment of a number of diseases, but does presently face some problems which limit the degree to which it is currently utilised. The major problem facing those who wish to use cryosurgery is effective control of the freezing zone. Clearly, it is important that the surgeon ensures that the whole of the affected tissue is cooled by a sufficient amount whilst avoiding, as far as possible, damage to adjoining tissue through inadvertent freezing.
The most common way for inducing the requisite cooling is to insert a plurality of metal cryoprobes into the vicinity of the tissue to be frozen. These are then cooled down to typically −200° C. using liquid nitrogen or adiabatic expansion of a highly pressurised gas such as argon. The tissue is cooled by thermal conduction from each of the cryoprobes and thus, in practice, an ice ball forms at the tips of the probes and grows in size to encompass the tissue to be frozen. The exact shape of the ice ball will clearly depend partly upon the distribution of the cryoprobes. Its shape will also depend on other factors such as the thermal conductivity of the tissue, the thermal diffusivity of the tissue and blood perfusion and the extent to which these parameters are anisotropic.
One way to control the growth of the ice ball is to generate an ultrasonic image of the treatment site and to use this to judge when the ice ball has grown to an appropriate size. The progress of the ice ball may be seen as a contrast between the frozen and unfrozen tissue. However, it will be appreciated that this will show only tissue which is at freezing point and therefore gives no information as to the temperature reached by tissue within the ice ball, i.e. whether it is has achieved the required −40° C.
Furthermore, the applicants have appreciated that there is a very high thermal inertia associated with the heat conduction and phase transition processes involved which means that significant skill is required on the part of the surgeon to remove the source of cooling agent sufficiently in advance of the ice ball actually reaching the boundary of healthy tissue to take account of its continued growth after the source of cooling has been removed.
It is possible to model the thermal behaviour of the tissue in order to predict where to place the cryoprobes to try to attain the desired temperature distribution and also to try to predict when to remove the source of cooling. An example of such a method is given in U.S. Pat. No. 5,647,868. Whilst being an apparently promising approach, it is beset with a number of problems and is not sufficiently accurate to be relied upon in real use.
Firstly, the thermal properties of bodily tissue tend to be highly inhomogeneous, thereby making the task of specifying the parameters extremely difficult and the resultant model extremely complicated. Furthermore, the thermal parameters are strongly temperature dependent which means that the model is highly non-linear. To give an example, the applicants have found experimentally that the thermal conductivity of bovine muscle frozen to about −20° C. is approximately three times larger than the corresponding value at 37° C. The thermal diffusivity of such tissue exhibits even larger changes—for example the diffusivity at −20° C. is almost four times larger than that at 37° C. Significant changes continue to occur as the tissue is cooled further down towards −200° C.
It is also found experimentally that there is no precisely defined freezing point in bodily tissue. Tissue starts to form ice crystals at −0.5° C. and the process continues to about −10° C. However, there is always approximately 10% of water within the tissue which never forms ice crystals. There is, therefore, no well-defined phase transition temperature as in the case of pure water.
In spite of these difficulties, the Applicants have developed a model in order to demonstrate some of the problems inherent in current cryosurgical techniques. In planar geometry, the propagation of the freezing zone may be modelled as follows. The tissue is assumed to be semi-infinite and homogenous. It is also assumed that the tissue is at a constant initial temperature. The distance, X, of the frozen front from the applicator at time t is given by the following expression:X=2λ√χ1t  (Eq 1)where χ1 is the thermal diffusivity of the frozen tissue (assumed to be temperature independent). The parameter λ characterises the properties of the freezing front—i.e. the interface between frozen and unfrozen tissue—and is therefore dependent on the thermal properties of the frozen and unfrozen tissue, the freezing point and latent heat of freezing (L) of the tissue and on the temperatures of the applicator and tissue respectively.
Table 1 gives some values for the parameter λ for various values of the applicator temperature. λ is calculated using realistic values for the various parameters, the values being given below Table 1. The symbols used for these parameters throughout are as follows:    χ=thermal diffusivity (m2/s)    K=thermal conductivity (W/mK)    L=latent heat (J/kg)    T=temperature
The table also gives the time taken in minutes for the planar freezing front to travel a distance of 2.5 mm, 5 mm and 10 mm respectively from the applicator for the various applicator temperatures.
TABLE 1ApplicatorDistanceDistanceDistancetemperatureParameter2.5 mm,5 mm,10 mm,in deg. C.λTime in min.Time in min.Time in min.−2000.7300.110.441.7−1500.6520.140.552.2−1000.5480.190.773.1−500.3940.371.56.0−400.3510.471.97.5−300.3010.642.610−200.2391.04.116−100.1552.49.639χfrozen = 4.5 · 10−7 m2/sχnonfrozen = 1.2 · 10−7 m2/sKfrozen = 1.5 W/mKKnonfrozen = 0.5 W/mKL = 2.5 · 105 J/m3 (80% water)Tfreeze = −0.5° C.Tnormal = 37° C.
As mentioned previously, the thermal diffusivity (χ) is significantly higher in frozen tissue than in unfrozen tissue. This phenomenon compensates for the tendency for the velocity of the freezing front to be reduced as a result of the required latent heat of freezing.
Substituting a cylindrical applicator which is assumed to be infinitely long, the results in Table 2 are achieved for various values for the power drained out per centimetre along the applicator.
TABLE 2Power drainedDistanceDistanceDistanceout, in W/cmParameter2.5 mm,5 mm,10 mm,lengthλTime in min.Time in min.Time in min.100.5150.20.93.580.4640.31.14.360.4020.41.45.740.3190.62.39.120.1981.55.924χfrozen = 4.5 · 10−7 m2/sχnonfrozen = 1.2 · 10−7 m2/sKfrozen = 1.5 W/mKKnonfrozen = 0.5 W/mKL = 2.5 · 105 J/m3 (80% water)Tfreeze = −0.5° C.Tnormal = 37° C.
When the applicator is removed, the situation may be modelled as a semi-infinite uniformly frozen tissue brought into contact with semi-infinite non-frozen tissue. Again, the distance, R, of the propagating frozen zone can be expressed by:R=2λ√χ1t  (Eq 2)
This time, however, λ is as given in Table 3. This also gives the time in minutes for the freezing front to continue to propagate a distance of 2.5 mm, 5 mm and 10 mm respectively.
TABLE 3InitialtemperatureDistanceDistanceDistanceof frozenParameter2.5 mm,5 mm,10 mm,tissue, in deg. C.λTime in minTime in minTime in min−2000.4480.31.14.6−1500.3640.41.77.0−1000.2550.93.6 (14)−500.1055.2(21) (84)−400.067(12.9)(52)(1608)−300.024(100)  (402) (6400)−24.20∞∞∞>−24.2NegativeThawingThawingThawingχfrozen = 4.5 · 10−7 m2/sχnonfrozen = 1.2 · 10−7 m2/sKfrozen = 1.5 W/mKKnonfrozen = 0.5 W/mKL = 2.5 · 105 J/m3 (80% water)Tfreeze = −0.5° C.Tnormal = 37° C.
From this it will be seen that there is a value of the initial temperature of the frozen tissue which gives a value for λ of 0 thereby representing a critical temperature for the frozen tissue, below which the freezing zone will continue to propagate after the applicator is removed, but above which the freezing zone will not continue to propagate. Of course, it will be appreciated that the critical value in this model is a higher temperature than that which is required to ablate malignant tissue effectively (i.e. approx −40° C.) and thus, conversely, if the present tissue is at a sufficiently low temperature, the freezing zone will continue to propagate.
Eventually the freezing zone will retreat as the tissue thaws again. The retreat of the freezing front will, however, be much slower than its advance. There are several reasons for this. Firstly, the thermal diffusivity in non-frozen tissue is, as mentioned above, significantly lower than in frozen tissue. Secondly, the transfer of latent heat of melting slows the process. Thirdly, the temperature difference between freezing point (e.g. −0.5° C.) and 37° C. is much less than the difference between the freezing point and an applicator at −200° C. In an example of the above, the velocity of propagation of the thawing zone in the case of a 40° C. applicator brought into contact with tissue frozen down to −200° C. would be only 7% of the velocity of the freezing boundary in the opposite direction when a −200° C. applicator is applied.
It should be borne in mind that the above models do not take blood perfusion into account. Blood perfusion also introduces an uncertain factor if included in a thermal model of tissue. During the freezing process, blood vessels are occluded and the blood circulation is shut down. After thawing, an edema develops in the outer margins of the frozen region and blood perfusion does not necessarily return to its original value. The perfusion in such margins can, therefore, be different during successive freezing periods.
Blood perfusion will tend to speed up the thawing process and the effect is likely to dominate at timescales longer than the blood perfusion relaxation time, typical values for which are 5–10 minutes in muscle tissue and 20–30 minutes in poorly perfused adipose tissue. The figures given in brackets in Table 3 are those which may be significantly affected by blood perfusion.
It will be seen from the above that there is a significant problem in the practical application of cryosurgery that thermal perturbations are highly diffused—i.e. smoothed out with time and distance from the source of the perturbation (e.g. the cryoapplicator). The problem can be reduced by minimising the distance between the applicator and the edge of the tissue which is to be frozen, but the extent to which this may be done is limited by the number of applicators required. Typically, the distance between an applicator and the edge of the affected tissue is of the order to 10 mm–15 mm.
It can be established from the crude models set out previously that in practical techniques the propagation of the freezing boundary towards the edge of the desired tissue cannot be stopped instantaneously—e.g. by cutting off the flow of cryogen thereto. There will be a time delay before any change of the applicator condition will have an influence on the edge of the freezing boundary.
It can be shown that in the above models this time delay is proportional to the square of the distance between the boundary and the cryoapplicator and inversely proportional to the thermal diffusivity of the frozen tissue. Thus, for a distance in the range of 10 mm–15 mm, the delay is of the order of 4–8 minutes. The freezing boundary will, of course, propagate several millimetres during this time. It is therefore necessary for the surgeon to judge when to adjust the cooling conditions significantly before the outer limit of the freezing zone is reached.
It is possible to protect some healthy body structures by passing heated water through them in order to prevent freezing thereof. In one particular example, water at approximately 37° C. may be passed through the urethra during cryosurgical ablation of the prostate. This solution is, however, limited by the maximum temperature which can be tolerated by the healthy tissue structure, which is only approximately 45° C. The rate of heat conduction into the tissue is therefore relatively low. Furthermore, the water is soon cooled and thus its temperature can differ significantly along the length of the region it is designed to protect. A discussion of this problem and an attempt at a solution thereof is given in U.S. Pat. No. 6,017,361. There remains, however, the limitation that this approach can only be used to protect structures in the vicinity of regions through which heated water can be passed.
The maximum temperature of such water which may be tolerated in the human body in order to prevent hyperthermia or protein denaturation is approximately 45° C. On the other hand, successful cryotherapy requires temperatures which are no greater than approximately −40° C. to −50° C. Significant amounts of power are required to maintain such a large temperature drop across only approximately 3 mm–5 mm of tissue. Typically, a power density of 1W–2W per square centimetre is required to maintain this thermal gradient. This would require a high flow rate and effective thermal transfer between the water and the tissue to be protected.
Where a critical structure within the body is protected by passing water either at body temperature or a slightly elevated temperature through a suitable tube the separation between the temperature source and the freezing boundary may be significantly smaller than between a cryoapplicator and the freezing zone, e.g. of the order of 3 mm–5 mm, but the time delay in such cases is still significant since the relevant diffusivity is that of non-frozen tissue which, as stated earlier, is significantly lower than that in frozen tissue. Thus, although the distance is reduced by a factor of three as compared to the cryoapplicator time delay, the heating time delay is only reduced by approximately 50% due to this reduced thermal diffusivity and so is of the order of 2–4 minutes.
It is an object of the present invention to provide an improved way of protecting surrounding body tissue during cryosurgery.